Shape memory bioresorbable polymer peripheral scaffolds

ABSTRACT

Bioabsorbable scaffolds having high crush recoverability, high fracture resistance, and reduced or no recoil due to self expanding properties at physiological conditions are disclosed. The scaffolds are made from a random copolymer of PLLA and a rubbery polymer such as polycaprolactone.

This application is a continuation of U.S. application Ser. No.15/588,514 filed May 5, 2017, which is a continuation of U.S.application Ser. No. 14/603,087 filed Jan. 22, 2015, now U.S. Pat. No.9,668,894 issued Jun. 6, 2017, which is a continuation of U.S.application Ser. No. 13/555,903 filed Jul. 23, 2012, now U.S. Pat. No.8,968,387 issued Mar. 3, 2015, all of which are incorporated in theirentireties by reference herein.

BACKGROUND OF THE INVENTION Field of the Invention

This invention relates polymeric medical devices, in particular,bioresorbable stents or stent scaffoldings.

Description of the State of the Art

This invention relates to radially expandable endoprostheses that areadapted to be implanted in a bodily lumen. An “endoprosthesis”corresponds to an artificial device that is placed inside the body. A“lumen” refers to a cavity of a tubular organ such as a blood vessel. Astent is an example of such an endoprosthesis. Stents are generallycylindrically shaped devices that function to hold open and sometimesexpand a segment of a blood vessel or other anatomical lumen such asurinary tracts and bile ducts. Stents are often used in the treatment ofatherosclerotic stenosis in blood vessels. “Stenosis” refers to anarrowing or constriction of a bodily passage or orifice. In suchtreatments, stents reinforce body vessels and prevent restenosisfollowing angioplasty in the vascular system. “Restenosis” refers to thereoccurrence of stenosis in a blood vessel or heart valve after it hasbeen treated (as by balloon angioplasty, stenting, or valvuloplasty)with apparent success.

Stents are typically composed of a scaffold or scaffolding that includesa pattern or network of interconnecting structural elements or struts,formed from wires, tubes, or sheets of material rolled into acylindrical shape. This scaffolding gets its name because it physicallyholds open and, if desired, expands the wall of the passageway.Typically, stents are capable of being compressed or crimped onto acatheter so that they can be delivered to and deployed at a treatmentsite.

Delivery includes inserting the stent through small lumens using acatheter and transporting it to the treatment site. Deployment includesexpanding the stent to a larger diameter once it is at the desiredlocation. Mechanical intervention with stents has reduced the rate ofrestenosis as compared to balloon angioplasty. Yet, restenosis remains asignificant problem. When restenosis does occur in the stented segment,its treatment can be challenging, as clinical options are more limitedthan for those lesions that were treated solely with a balloon.

Stents are used not only for mechanical intervention but also asvehicles for providing biological therapy. Biological therapy usesmedicated stents to locally administer a therapeutic substance. Amedicated stent may be fabricated by coating the surface of either ametallic or polymeric scaffold with a polymeric carrier that includes anactive or bioactive agent or drug. Polymeric scaffolds may also serve asa carrier of an active agent or drug. An active agent or drug may alsobe included on a scaffold without being incorporated into a polymericcarrier.

The stent must be able to satisfy a number of mechanical requirements.The stent must be capable of withstanding the structural loads, namelyradial compressive forces, imposed on the scaffold as it supports thewalls of a vessel. Therefore, a stent must possess adequate radialstrength. Radial strength, which is the ability of a stent to resistradial compressive forces, relates to a stent's radial yield strengthand radial stiffness around a circumferential direction of the stent. Astent's “radial yield strength” or “radial strength” (for purposes ofthis application) may be understood as the compressive loading, which ifexceeded, creates a yield stress condition resulting in the stentdiameter not returning to its unloaded diameter, i.e., there isirrecoverable deformation of the stent. When the radial yield strengthis exceeded the stent is expected to yield more severely and only aminimal force is required to cause major deformation. Radial strength ismeasured either by applying a compressive load to a stent between flatplates or by applying an inwardly-directed radial load to the stent.

Once expanded, the stent must adequately maintain its size and shapethroughout its service life despite the various forces that may come tobear on it, including the cyclic loading induced by the beating heart.For example, a radially directed force may tend to cause a stent torecoil inward. In addition, the stent must possess sufficientflexibility to allow for crimping, expansion, and cyclic loading.

Some treatments with stents require its presence for only a limitedperiod of time. Once treatment is complete, which may include structuraltissue support and/or drug delivery, it may be desirable for the stentto be removed or disappear from the treatment location. One way ofhaving a stent disappear may be by fabricating a stent in whole or inpart from materials that erodes or disintegrate through exposure toconditions within the body. Stents fabricated from biodegradable,bioabsorbable, bioresorbable, and/or bioerodable materials such asbioabsorbable polymers can be designed to completely erode only afterthe clinical need for them has ended.

The development of a bioresorbable stent or scaffold could obviate thepermanent metal implant in vessel, allow late expansive luminal andvessel remodeling, and leave only healed native vessel tissue after thefull absorption of the scaffold. A fully bioresorbable stent can reduceor eliminate the risk of potential long-term complications and of latethrombosis, facilitate non-invasive diagnostic MRI/CT imaging, allowrestoration of normal vasomotion, and provide the potential for plaqueregression.

To treat peripheral vascular disease percutaneously in the lower limbsis a challenge with current technologies. Long term results aresub-optimal due to chronic injury caused by the constant motions of thevessel and the implant as part of every day life situations. To reducethe chronic injury a bioresorbable scaffold for the superficial femoralartery (SFA) and/or the popliteal artery can be used so that thescaffold disappears before it causes any significant long term damage.However, one of the challenges with the development of a femoralscaffold and especially a longer length scaffold (4-25 cm) to be exposedto the distal femoral artery and potentially the popliteal artery is thepresence of fatigue motions that may lead to chronic recoil and strutfractures especially in the superficial femoral artery, prior to theintended bioresorption time especially when implanted in the superficialfemoral artery.

A scaffold in the SFA and/or the popliteal artery is subjected tovarious non-pulsatile forces, such as radial compression, torsion,flexion, and axial extension and compression. These forces place a highdemand on the scaffold mechanical performance and can make the scaffoldmore susceptible to fracture than less demanding anatomies. In additionto high radial strength, stents or scaffolds for peripheral vessels suchas the SFA, require a high degree of crush recovery. The term “crushrecovery” is used to describe how the scaffold recovers from a pinch orcrush load, while the term “crush resistance” is used to describe theforce required to cause a permanent deformation of a scaffold.

Therefore, an important goal for treatment of the SFA and/or thepopliteal artery is the development of bioabsorbable scaffold with highradial strength, high crush recovery, and high resistance to fracture orhigh toughness.

INCORPORATION BY REFERENCE

All publications and patent applications mentioned in this specificationare herein incorporated by reference to the same extent as if eachindividual publication or patent application was specifically andindividually indicated to be incorporated by reference, and as if eachsaid individual publication or patent application was fully set forth,including any figures, herein.

SUMMARY OF THE INVENTION

Embodiments of the present invention include a stent comprising ascaffold formed from a polymer tube configured for being crimped to aballoon, the scaffold having a pattern of interconnected struts and thescaffold having an expanded diameter when expanded from a crimped stateby the balloon, wherein the scaffold attains greater than about 80% ofits diameter after being crushed to at least 50% of its expandeddiameter; wherein the scaffold has a radial stiffness greater than 0.3N/mm and wherein the scaffold is made from a shape memory randomcopolymer of poly(L-lactide) (PLLA) and a rubbery polymer that is 0.1 to10 wt % or molar % of the copolymer, wherein the scaffold exhibits selfexpanding properties at 37 deg C. in physiological conditions.

BRIEF DESCRIPTION OF THE DRAWINGS

FIG. 1 depicts the Tg onset of four samples of PLLA andpoly(L-lactide-co-caprolactone) (PLLA-PCL) copolymers at three molarcompositions, 95:5, 90:10, and 85:15, as function of soak time in water.

FIG. 2 depicts a first embodiment of a crush recoverable scaffoldpattern.

FIG. 3 is a partial perspective view of a scaffold structure.

FIG. 4 depicts a second embodiment of a crush-recoverable scaffoldstructure.

FIG. 5A depicts aspects of the repeating pattern of closed cell elementsassociated with the pattern of FIG. 4.

FIG. 5B depicts aspects of the repeating pattern of closed cell elementsassociated with the pattern of FIG. 2.

FIGS. 6A and 6B are tables showing examples of scaffold features inaccordance with aspects of the disclosure.

FIG. 7 depicts a third embodiment of a crush-recoverable scaffoldstructure.

FIG. 8 depicts aspects of the repeating pattern of closed cell elementsassociated with the pattern of FIG. 7.

FIG. 9 shows the radial strength and stiffness for scaffold samples.

FIG. 10 shows the diameter of a 90:10 PLLA-PCL scaffold as a function oftime after expansion.

FIG. 11 depicts crush recovery results for 95:5 PLLA-PCL and 90:10PLLA-PCL scaffolds.

FIG. 12 shows discontinuity count in ring and connector links afteraxial fatigue testing for the PLLA, 95:5 PLLA-PCL, and 90:10 PLLA-PCLscaffolds.

FIG. 13 shows discontinuity percentage in ring and connector links afteraxial fatigue testing for the PLLA, 95:5 PLLA-PCL, and 90:10 PLLA-PCLscaffolds.

FIGS. 14 and 15 depict Finescan images of the 95:5 and the 90:10PLLA-PCL scaffolds, respectively, post deployment.

DETAILED DESCRIPTION OF THE INVENTION

The embodiments described herein are generally applicable to polymericimplantable medical devices, especially those that have load bearingportions when in use or have portions that undergo deformation duringuse. In particular, the methods can be applied to tubular implantablemedical devices such as self-expandable stents, balloon-expandablestents, and stent-grafts.

A stent or scaffold may include a tubular scaffold structure that iscomposed of a plurality of ring struts and link struts. The ring strutsform a plurality of cylindrical rings arranged about the cylindricalaxis. The rings are connected by the link struts. The scaffold comprisesan open framework of struts and links that define a generally tubularbody with gaps in the body defined by the rings and struts. Athin-walled cylindrical tube of may be formed into this open frameworkof struts and links described by a laser cutting device that cuts such apattern into the thin-walled tube that may initially have no gaps in thetube wall.

The scaffold may also be fabricated from a sheet by rolling and bondingthe sheet to form the tube. A tube or sheet can be formed by extrusionor injection molding. The scaffold can then be crimped on to a balloonor catheter for delivery into a bodily lumen.

A stent or scaffold of the present invention can be made partially orcompletely from a biodegradable, bioresorbable, and bioabsorbablepolymer. The stent can also be made in part of a biostable polymer. Apolymer for use in fabricating stent can be biostable, bioresorbable,bioabsorbable, biodegradable, or bioerodable. Biostable refers topolymers that are not biodegradable. The terms biodegradable,bioresorbable, bioabsorbable, and bioerodable are used interchangeablyand refer to polymers that are capable of being completely degradedand/or eroded into different degrees of molecular levels when exposed tobodily fluids such as blood and can be gradually resorbed, absorbed,and/or eliminated by the body. The processes of breaking down andabsorption of the polymer can be caused by, for example, hydrolysis andmetabolic processes.

Bioresorbable stents or scaffolds can be useful for treatment of varioustypes of bodily lumens including the coronary artery, superficialfemoral artery, popliteal artery, neural vessels, and the sinuses. Ingeneral, these treatments require the stent to provide mechanicalsupport to the vessel for a period of time and then desirably to absorbaway and disappear from the implant site. The important properties of abioresorbable stent or scaffolding include mechanical and degradationproperties. The mechanical requirements include high radial strength,high radial stiffness, minimal recoil, and high fracture toughness. Thedegradation properties include the absorption profile, for example, thechange in molecular weight, radial strength, radial stiffness, and masswith time. Specific aspects of the absorption profile include the timethat the stent maintains radial strength and radial stiffness, beforestarting to decrease and the total absorption time or absorption time(complete mass loss from implant site).

A stent scaffolding made from a bioresorbable polymer may be designed tomaintain its radial strength and/or radial stiffness once implanted toprovide mechanical support to the vessel for a prescribed time periodand maintain patency of the lumen. The radial strength must besufficiently high initially to support the lumen at a desired diameter.The period of time that the scaffold is required or desired to maintainpatency depends on the type of treatment, for coronary treatment it isabout 3 months. After this time period, the vessel is healedsufficiently to maintain an expanded diameter without support.Therefore, after this time period, the scaffolding may start to loseradial strength and/or radial stiffness due to molecular weightdegradation. As the scaffolding degrades further, it starts to losemechanical integrity and then experiences mass loss and eventuallyabsorbs away completely or there are negligible traces left behind. Thedecrease in radial strength and radial stiffness can also be the resultof mechanical fractures in certain anatomies where this is desired, orit could also be the result of a combination of the mechanical andchemical degradation.

Ideally, it is desired that once the stent support is no longer neededby the lumen, the bioresorbable scaffold should be resorbed as fast aspossible while also meeting all basic safety requirements during itsdegradation period. Such safety requirements can include a gradualdisintegration and resorption that does not allow release of fragmentsthat could cause adverse events such as thrombosis. In this way, thestent scaffold enables the vessel healing as well as enabling theadvantages mentioned herein of a bioresorbable scaffold to the greatestextent. It is desirable for a bioresorbable scaffold to have anabsorption time of about 18 to 26 months for coronary vascularapplication, of about 12-26 months for a peripheral application (e.g.,superficial femoral artery (SFA), tibial, and/or politeal artery), 18-24months for neural applications, and less than a year for nasalapplications.

With respect to radial strength and stiffness, a stent should havesufficient radial strength and/or stiffness to withstand structuralloads, namely radial compressive forces, imposed on the stent so thatthe stent can supports the walls of a vessel at a selected targetdiameter for a desired time period. A polymeric stent with adequateradial strength and/or stiffness enables the stent to maintain a lumenat a desired diameter for a sufficient period of time after implantationinto a vessel.

In addition, the stent should possess sufficient toughness or resistanceto fracture to allow for crimping, expansion, and cyclic loading withoutfracture or cracking that would compromise the function of the stent.The toughness or resistance to fracture of the scaffold material can becharacterized for a material by the elongation at break and for a stentby the number and degree of cracks in a scaffold during use, such asafter crimping or deployment or dilation to a target diameter. Theseaspects of the use of the scaffold involve deformation of various hingeportions of the structural elements of the scaffold.

Some bioresorbable polymers, for example, semi-crystalline polymers, arestiff or rigid under physiological conditions within a human body andhave been shown to be promising for use as a scaffold material.Specifically, polymers that have a glass transition temperature (Tg)sufficiently above human body temperature which is approximately 37° C.,should be stiff or rigid upon implantation. Poly(L-lactide) (PLLA) isattractive as a stent material due to its relatively high strength and arigidity at human body temperature, about 37° C. As shown in Table 1,PLLA has high strength and tensile modulus compared to otherbiodegradable polymers. Since it has a glass transition temperature wellabove human body temperature, it remains stiff and rigid at human bodytemperature. This property facilitates the ability of a PLLA stentscaffolding to maintain a lumen at or near a deployed diameter.

Other rigid bioresorbable polymers include poly(D-lactide) (PDLA),polyglycolide (PGA), and poly(L-lactide-co-glycolide) (PLGA). The PLGAinclude those having a mole % of (LA:GA) of 85:15 (or a range of 82:18to 88:12), 95:5 (or a range of 93:7 to 97:3), or commercially availablePLGA products identified being 85:15 or 95:5 PLGA. Rigid polymers mayrefer to polymers that have a Tg higher than human body temperature orwithin 5 deg C. of human body temperature.

TABLE 1 Comparison of properties of bioresorbable polymers. Elonga-Tensile Tensile tion at Tm Tg Strength Modulus break Absorp- (° C.) (°C.) (MPa) (MPa) (%) tion Rate PLLA 175  65 28-50 1200-2700  6 1.5-5years P4HB  60 −51 50  70 1000  8-52 weeks PCL  57 −62 16 400 80 2 yearsPDO  110¹  −10¹ 1.5^(1,2)   30²  35³ 6-12¹ 6² PGA 225  35 70 6900  <3 6weeks DL- Amorphous 50-53 16 400 80 2 PLA years P3HB 180  1 36 2500   32 years PLLA (poly(L-lactide); P4HB (poly-4-hyroxybutyrate); PCL(polycaprolactone); PGA (polyglycolide); DL-PLA (poly(DL-lactide); P3HB(poly-3-hydroxybutyrate); PDO (p-polydioxanone) All except PDO, Martinet al., Biochemical Engineering 16 (2003) 97-105. ¹Medical Plastics andBiomaterials Magazine, March 1998. ²Medical Device Manufacturing &Technology 2005. ³The Biomedical Engineering Handbook, Joseph D.Bronzino, Ed. CRC Press in Cooperation with IEEE Press, Boca Raton, FL,1995.

The strength, stiffness, and the fracture toughness of such polymers canbe improved through various processing methods (e.g., radial expansionand suitable choice of associated processing parameters). However, thereis still strong incentive to improve upon polymers such as PLLA asscaffold materials not only for coronary applications, but to tailor itfor various peripheral applications as well. In particular, suchpolymers may be improved upon to reduce chronic recoil inward from adeployed diameter and to reduce strut fractures due to fatigue motionsimposed on scaffolds implanted in such vessels as the popliteal arteryand in the superficial femoral artery.

Embodiments of the present invention include bioresorbable scaffoldswith reduced risk of scaffold fracture and improved recoil property ascompared to PLLA scaffolds. Embodiments include scaffolds composed ofcopolymers that exhibit shape memory properties and are self-reinforcingonce implanted in vessel of a human patient. Shape-memory propertiesrefer to the ability of a material to return from a deformed state(temporary shape) to an original (permanent) shape induced by anexternal stimulus (trigger), such as a temperature change or thermaltransition. A scaffold of the present invention has a tendency return orself expand outward towards a fabricated diameter when deployed fromcrimped or reduced diameter.

The temperature of the thermal transition of a shape memory polymer isreferred to as a transition temperature (Ttrans) which is thetemperature around which a material changes from one state to another.In general, a Ttrans can be either a melting temperature (Tm) or glasstransition temperature (Tg). In the present invention, the Tg of theshape memory polymer of the scaffold is the relevant thermal transitionthat induces shape memory behavior of an implanted scaffold. Ttrans istypically determined by differential scanning calorimetry (DSC),thermomechanical analysis (TMA) or dynamic mechanical thermal analysis(DMA). DSC measures the change in heat capacity, TMA measures the changein coefficient of thermal expansion, while DMA measures the change inelastic modulus during the thermal transition. Due to intrinsicpolydispersity in molecular weights and imperfect spatial distributionof network chains, the unique thermal characteristics of a polymershould be defined as a temperature range rather than at one specifictemperature. For the ease of comparison, however, a single Ttrans (Tm orTg) value taken from the peak or midpoint of a broader transition isoften reported in literature.

For a particular shape memory copolymer with specific chemicalcomponents and their composition (e.g., 95:5 PLLA-PCL), the broadness ofa thermal transition may be modified in several ways. As indicatedabove, the broadness of Tg is a function of polydispersity in molecularweights. The polydispersity index (PDI) or heterogeneity index of apolymer, is a measure of the distribution of molecular mass in a givenpolymer sample. The PDI is calculated as the weight average molecularweight, Mw, divided by the number average molecular weight, Mn. Thus,the PDI can be manipulated to achieve a selected broadness of thethermal transition.

The width of thermal transition of an SMP is also dependent on thecrystallinity. Crystalline segments tend to exhibit a sharp transitionwith a relatively narrow temperature range while amorphous segments tendto display a glass transition range tens of degrees wide. Therefore,crystallinity of the scaffold can be manipulated through processing toachieve a selected broadness of the transition. For example, thecrystallinity of a particular copolymer can be adjusted throughannealing which increases crystallinity.

The copolymer of the scaffold is mostly poly(L-lactide) and a smallamount of rubbery polymer. The small amount of rubbery polymer togetherwith the poly(L-lactide) exhibit shape memory behavior upon deploymentin a vessel at physiological conditions. In addition, the rubberypolymer increases the fracture resistance of the scaffold which preventsor reduces fractures when the scaffold diameter changes, for example,when the scaffold is crimped from a fabricated diameter to a reducedcrimped diameter or when the scaffold is expanded from the crimpeddiameter to a target diameter during deployment.

Physiological conditions include a temperature at or about 37 deg C. andexposure to bodily fluids, in particular, a wet or aqueous environment.A rubbery polymer refers to a polymer that is more flexible, and thus,has a lower modulus and/or greater elongation at break than a rigidpolymer such as PLLA. For example, a rubbery polymer may have a tensilemodulus less than 1000 MPa or less than 500 MPa and/or an elongation atbreak greater than 10% or greater than 50%.

The copolymer scaffolds are fabricated from copolymer tubing at a givendiameter and crimped down to a reduced diameter over a catheter balloon.The crimped scaffold is deployed from the reduced diameter to a targetor nominal expansion diameter in a vessel. The scaffold may befabricated so that it exhibits an outward force at the target diameterby oversizing the scaffold (i.e., as-fabricated diameter greater thanthe target diameter) relative to the target diameter. The outward forceis a residual force in excess of a radial outward force that maintainsthe lumen at the target diameter. The oversizing is advantageous due tothe increased flexibility or lower modulus of the shape memory andreinforcing scaffold.

The oversized scaffold is made from a tube with a diameter greater thanthe target diameter. The tube may be oversized by a tubing expansiongreater than the intended target diameter. The expansion also results insufficient scaffold radial strength and stiffness to treat a stenosedartery at early duration of the implant. Due to the oversizing, thedeployed scaffold applies the outward force on the vessels walls at thetarget diameter or a diameter slightly greater than the target diameter.The outward force is analogous to the chronic outward force applied byself-expanding stents such as nitinol stents.

The self-expanding or shape memory scaffolds of the present inventionprovide sufficient radial strength, radial stiffness, and recoil toprovide patency to or support a vessel and at a target diameter.Additionally, the scaffolds of the present invention have a greaterresistance to fracture compared to a PLLA scaffold.

At physiological conditions of 37 deg C. and hydration in blood, thebioresorbable copolymer exhibits a thermal transition which results inshape memory behavior that includes a tendency for the scaffold toself-expand, to exert an outward residual force on the vessel wall, orboth. The shape memory behavior, however, acts in a time-dependentfashion so that the mechanical properties of the scaffolds such asradial strength and compression recovery are enhanced with time uponimplantation. In contrast to permanent nitinol stents, theself-reinforcing bioresorbable scaffold will degrade, reducing itsresidual outward force with time, transform from independentload-bearing member into tissue-incorporated composite, and ultimatelydisappear without causing any significant chronic vascular injury.

The mechanism of controlled reduction in scaffolding property (radialstrength/stiffness/strain recovery) will reduce tissue compliance(vascular compliance in case of a vascular implant) mismatch of theimplanted segment. Specifically, the scaffold undergoes degradation inthree phases.

In the first phase, the molecular weight is reduced primarily due tohydrolytic degradation of the polymer chains. During this first phase,the mechanical properties such as radial strength and radial stiffnesschange very little or not at all so that the scaffold supports thevessel at the target diameter during this phase. This allows vessel wallremodeling that will enable the vessel to support itself at theremodeled diameter once the scaffold is gone. In the second phase,mechanical properties including the radial strength and the residualoutward force of the scaffold decrease. The decrease in radial strengthin particular results in the transfer of the load or support of thevessel wall from the scaffold to the vessel wall. Also, in the secondphase, the scaffold begins to break up due to the deterioration ofmechanical properties. Prior to breaking up, the scaffold is preferablycovered by tissue. In the third phase, which may overlap the secondphase, the scaffold erodes, eventually completely away from the implantsite. It is during the second and third phases that the vesselcompliance increases gradually from that of the scaffold to that of thenatural compliance of an un-scaffolded vessel at a remodeled, increaseddiameter. The first and second phases could also be in the order ofmechanical degradation first followed by a chemical degradation.

The residual outward force of the scaffold is lower than that of thechronic outward force of a self expanding nitinol scaffold and willtherefore create less arterial injury than the nitinol implant. Moreoverthe axial and radial flexibility provided by the disclosed pattern ofthe scaffold will reduce even the acute injury that would otherwisepotentially occur even before any of either the chemical or mechanicaldegradation starts.

A PLLA scaffold has a Tg of about 60-65 deg C., and therefore, tends notto exhibit substantial shape memory properties upon expansion in avessel. Specifically, the PLLA scaffold is expanded to a targetdiameter, but does not exhibit a large tendency to self expand atphysiological conditions and also will not apply a residual force inexcess of that for maintaining patency of the vessel. The largedifference between the Tg of the scaffold and physiological conditions(e.g., 37 deg C. and wet) is responsible for the absence of a thermaltransition that would provide the tendency to self expand or a residualforce.

The rubbery polymer has a Tg in dry or wet conditions less than PLLA indry or wet conditions, respectively. Therefore, the copolymer of thescaffold of the present invention has a glass transition temperature(Tg) that is lower than that of the PLLA scaffold. The reduction in theTg of the scaffold closer to or below the physiological conditions of 37deg C. contributes to the shape memory property upon expansion in avessel to a target diameter. The scaffold copolymer Tg is sufficientlyclose to the physiological temperature for the scaffold to exhibit atendency to self expand towards an original fabricated diameter. Thistendency reduces or eliminates recoil of the scaffold. If the scaffoldis oversized, then the scaffold applies a residual outward force on thevessel wall in excess of that required to maintain the target diameterof the vessel. The residual force further reduces or eliminates recoilof the scaffold.

The Tg of the copolymer of the scaffold in dry or wet conditions may be5 to 10 deg C., 10 to 20 deg C., 10 to 30 deg C., 20 to 30 deg C., lessthan the Tg of the PLLA scaffold. Alternatively, the Tg of the copolymerscaffold may be 30 to 37 deg C., 37 to 50 deg C., 37 to 55 deg C., 37 to40 deg C., 40 to 45 deg C., 45 to 50 deg C., or 50 to 55 deg C. Therubbery polymer, as a homopolymer, may have a Tg less than 37 deg C.,less than 25 deg C., 0 to 25 deg C., or −70 to 0 deg C.

The inventors have observed that the Tg of the copolymer may decreasewhen exposed to moisture or hydrated such as by soaking in water orimplanted in a physiological environment. The decrease in Tg may beattributed to the water plasticizing the polymer. Plasticizing refersgenerally to increasing the plasticity of a material, where plasticityor plastic deformation describes the deformation of a polymer undergoingnon-reversible changes of shape in response to applied forces. Withoutbeing limited by theory, a plasticizer, in this case water, works byembedding itself between the chains of a polymer, spacing them apart(increasing the “free volume”), and thus lowering the glass transitiontemperature and making it softer. As a result of plasticizing, themechanical properties of the copolymer and scaffold change, inparticular, modulus of the copolymer decreases which decreases theradial stiffness. The inventors have found that the stability of Tg overtime and the degree of decrease in Tg is a function of the compositionof the rubbery polymer in the copolymer.

In some embodiments, the Tg of the copolymer is both stable for aselected period of time and exhibits a thermal transition providingshape memory properties to the scaffold when it is exposed to an aqueousenvironment or bodily fluids. The selected period of stability may be upto 4 days, up to 1 month, or up to 3 months after implantation. Forexample, the Tg of the copolymer may change by less than 15% after it ishydrated, either in water or in a physiological environment during suchtime. A stable Tg may be preferred so that the radial stiffness of thescaffold remains stable while the vessel wall is remodeling.

The Tg of a hydrated copolymer can remain stable or change by less than15%, or more narrowly, by between 2 to 5% when the copolymer has lessthan a selected weight or mole percent of the rubbery polymer. Above aselected composition of the rubbery polymer, the Tg of the copolymer canchange significantly over time when hydrated. FIG. 1 depicts the Tgonset of four samples of PLLA and poly(L-lactide-co-caprolactone)(PLLA-PCL) copolymers at three molar compositions, 95:5, 90:10, and85:15, as function of soak time in water. As shown in FIG. 1, the Tg ofthe PLLA and 95:5 PLLA-PCL is stable over 100 hr soaking. The Tg's ofthe 90:10 and 85:15 PLLA-PCL, however, change significantly over the 100hr soak time.

In some embodiments, the rubbery polymer may be 0.5 to 1%, 1 to 5%, 2 to5%, 3-5% (by weight or molar) of the copolymer of the scaffold. In otherembodiments, the rubbery polymer may be 5 to 15%, 5 to 15%, or 15 to 25%of the copolymer of the scaffold.

In general, an increase in the amount of rubbery polymer increasesfracture toughness of the copolymer and the shape memory properties,which provides increased resistance to fracture of the scaffold whencrimped or deployed and reduction in recoil. However, the increase inrubbery polymer may decrease the strength and modulus of the copolymer,which can reduce radial strength and radial stiffness of the scaffold.Therefore, the amount of rubbery polymer should not be too high suchthat the resulting scaffold does not possess sufficient radial strengthor stiffness to scaffold a stenosed artery.

Many rubbery polymers that could be used to provide shape memoryproperties have a lower strength than PLLA. Therefore, the radialstrength of the scaffold is expected to decrease as the amount of therubbery polymer increases. For example, as shown in Table 1,polycaprolactone has a lower tensile strength than PLLA. It is preferredthat the composition of the rubbery polymer in the copolymer is lessthan or equal to 5 or 15% (weight or molar percent) relative to PLLA ofthe scaffold.

However, in some treatments, the scaffold is not required to support avessel at an increased diameter, so radial strength is less or notimportant. In such treatments, the scaffold may serve as a sustaineddrug delivery scaffold. For such applications, a flexible copolymerscaffold may be used with greater than 15% rubbery polymer, for example,15 to 30%, 15 to 50%, or 50 to 75% rubbery polymer.

The copolymer of the scaffolds may exhibit phase separation of therubbery polymer from the PLLA. Thus, the polymeric scaffold made of thecopolymer may be composed of soft domains and hard domains. The softdomains may be dispersed throughout the hard domains. The hard domainsare partially crystalline and the soft domains are mostly amorphous orcompletely amorphous. The thermal transition temperatures (either Tg orTm in a physiological environment such as blood at 37 deg C.) of thesoft domain may be at 37 deg C., between 37 and 47 deg C., or anytemperature below 37 deg C.

The characteristic length (diameter, length, width) of the domains canbe 250 nm to 5 um, or more narrowly 250 nm to 1000 nm, 100 nm to 500 nm,500 nm to 1 μm, or greater than 1 μm.

The copolymer thermal transition for hard domains (either glasstransition or melt transition in aqueous environment such as blood) maybe at a temperature greater than the physiologic temperature of 37 degC. to preserve the desired scaffold deployment dimensions and preventsubsequent creep and relaxation properties. These deployment dimensionsmay be provided to the copolymer scaffold during melt extrusion of thetubing and pressurized tubing expansion (described below) at atemperature greater than the Tg, but less than the Tm.

Examples of biodegradable rubbery polymers include but are not limitedto polyhydroxyalkanoates (PHA), poly(4-hydroxybutyrate) (P4HB),poly(ε-caprolactone), (PCL) poly(trimethylene carbonate) (PTMC),poly(butylene succinate) (PBS), poly(p-dioxanone) (PDO), poly(esteramides) (PEA), and biodegradable polyurethanes. A preferred type ofcopolymer is a random copolymer of PLLA and the rubbery polymer.

The molecular weight of the shape memory copolymer of the scaffoldpost-sterilization may be 60 to 150 kDA, or more narrowly 80 to 100 kDa.The molecular weight of the shape memory copolymer of the scaffoldpre-sterilization may be 100 to 400 kDA, or more narrowly 150 to 300kDa.

Embodiments can also include block copolymers of the PLLA and therubbery polymer. These block copolymers can include linear blockcopolymers such as diblock (AB), triblock (ABA), or, generally,multiblock copolymers (ABABA) and star block copolymers. Additionalblock copolymers include hyperbranched-like polymers, comb-likepolymers, dendrimer-like star polymers, and dendrimers.

Embodiments of the present include scaffolds made out the copolymerwhich is not blended with another polymer, except for small amounts ofanother polymer, e.g., less than 1 wt % or incidental mixing of thecopolymer with a coating at the coating/copolymer interface.

Soft domains may be modified by post plasticizing the domains withplasticizers that swell, but do not dissolve the copolymer. Theplasticizing will further reduce the Tg of the polymer. For example, aPLLA-PCL copolymer scaffold can be plasticized with acetone orchloroform.

As indicated above, the thermal transition of the shape memory copolymercan occur around, not necessarily only at a thermal transition Trans,for example, Tg. Therefore, the shape memory copolymer of the scaffoldneed not have a Tg at this temperature for the scaffold to exhibit shapememory properties. The thermal transition may have varying degrees ofbroadness, i.e., the thermal transition may occur within a certain ΔTbelow the Tg of the copolymer. The copolymer can be designed to have aselected broadness of both Tg and Tm. For example, a copolymer may behave a broad Tg and a sharp Tm for a shape memory scaffold with a Tg>37deg C. A broad Tg or Tm may be defined as a ΔT of between 5 and 20 degC. A narrow Tg or Tm may be defined as a ΔT of between 5 and 10 deg C.Alternatively, a copolymer may have a sharp Tm and sharp Tg for a shapememory scaffold with Tg<37 deg C.

A shape memory polymer may also exhibit more than one Tg which may occurin block copolymers or random copolymers with a high composition ofrubbery polymer, e.g., greater than 15 weight percent of rubberypolymer. A sharp Tm and two independent Tg's may provide additionaldegrees of freedom to control the shape memory evolution of the implant.

A preferred embodiment is a scaffold made completely of the shape memorycopolymers. In some embodiments, the shape memory copolymer is notblended with another polymer. The scaffold, however, may include smallamounts of other additives such as antioxidants, inorganic reinforcingagents, or therapeutic agents, for example, 0.1 to 2 wt %.

Another embodiment is a blend of the copolymer and another polymer. Thescaffold may further include a polymer and drug coating.

In an alternate embodiment, a PLLA scaffold may have a coating of theshape memory copolymer. The PLLA scaffold can be dip or spray coatedwith the shape memory polymer, for example PLLA-PCL. The coating mayhave a thickness of 2 to 10 microns or 10 to 20 microns. The shapememory coating may reduce or prevent PLLA crack propagation and strutseparation prior to tissue coverage. The shape memory polymer may alsobe applied to widen the overstretch window prior to fracture for ease ofphysician utilization of the scaffolds.

Another advantage of the shape memory scaffold of the present inventionis that physical aging that causes embrittlement of the scaffold polymeroccurs over a shorter time frame than for a PLLA scaffold. In general,after fabrication, a polymer of a scaffold undergoes a process calleddensification that occurs over a period of time until the propertiesstabilize. The time it takes for the polymer to stabilize is higher forhigher Tg's. At the lower Tg of the shape memory copolymer, the physicalaging process is faster and stabilizes faster. For example, for a Tgbetween 40 and 50 deg C., the polymer is expected to be stabilizedwithin 2 weeks when it is stored at RT, while it might take more than 1month to stabilize at RT for polymer with Tg at about 60 deg C. As aresult, the products used for implantation will have high consistency ofproperties.

Additionally, the crystallinity of the shape memory random copolymer ofPLLA and rubbery polymer is lower than PLLA. The rubbery polymer unitsincorporated in the copolymer chains tend to disrupt the formation ofcrystalline segments. The crystallinity of the PLLA scaffold may be 45to 55%. In contrast, the crystallinity of the shape memory copolymerscaffold may be 10 to 40% or 20 to 30%. The decreased crystallinityincreases the degradation rate, as compared to a PLLA scaffold withhigher crystallinity.

Stent scaffold patterns made from PLLA for SFA and poplitealapplications have been designed which have high crush recovery and crushresistance. Crush recovery describes the recovery of a scaffoldsubjected to a pinch or crush load. Specifically, the crush recovery canbe described as the percent recovery to the scaffold pre-crush shape ordiameter from a certain percent crushed shape or diameter. Crushresistance is the minimum force required to cause a permanentdeformation of a scaffold.

The crush recoverable PLLA scaffolds are disclosed in US2011/0190872,US2011/0190871, and U.S. patent application Ser. No. 13/549,366. Thesescaffolds attain greater than about 80% of its diameter after beingcrushed to at least 50% of its expanded diameter. The scaffolds alsohave a radial stiffness greater than 0.3 N/mm. Such scaffolds also havea normalized radial strength of at least 0.4 N/mm. The inventors havefound a scaffold for use in peripheral application should have theseminimum values of radial stiffness and radial strength.

Embodiments of the present invention further include bioresorbablescaffolds that in addition to high crush recovery, the selected radialstiffness and radial strength, also possess high resistance to fractureand self expanding properties that reduce or eliminate recoil upondeployment. These embodiments include the previously disclosed crushrecoverable scaffolds and additional crush recoverable scaffoldsdisclosed herein made from the shape memory copolymer disclosed herein.The inventors have found that a suitable combination of scaffold andshape memory copolymer can result in a scaffold having desiredperformance characteristics.

The scaffolds of the present invention for peripheral (SFA) applicationsusually have lengths of between about 36 and 40 mm or even between 40and 200 mm when implanted in the superficial femoral artery, as anexample. The scaffold for the SFA applications may have a pre-crimpingdiameter of between 5-10 mm, or more narrowly, 6-8 mm. The scaffold forSFA may have a wall thickness of about 0.008″ to 0.014″ and configuredfor being deployed by a non-compliant or semi-compliant balloon, e.g.,6.5 mm diameter, from about a 1.8 to 2.2 mm diameter (e.g., 2 mm)crimped profile. The SFA scaffold may be deployed to a diameter ofbetween about 4 mm and 10 mm, or more narrowly, 7 to 9 mm.

These crush recoverable copolymer scaffolds can attain greater thanabout 80% of their diameter after being crushed to at least 50% of itsexpanded diameter. Additionally, such crush recoverable copolymerscaffolds have a normalized radial strength, as measured by techniquesdescribed herein and in cited applications, of greater than about 0.3N/mm, or between about 0.3 and 1.2 N/mm or between about 0.3 and 1.2N/mm, and a radial stiffness of greater than about 0.3 N/mm or betweenabout 0.3 and 2 N/mm.

A first embodiment of a crush recoverable scaffold pattern is depictedin FIG. 2. FIG. 2 depicts the pattern 200 which includeslongitudinally-spaced rings 212 formed by struts 230. The pattern 200 ofFIG. 2, represents a tubular scaffold structure (as partially shown inthree dimensional space in FIG. 3), so that an axis A-A is parallel tothe central or longitudinal axis of the scaffold. FIG. 3 shows thescaffold in a state prior to crimping or after deployment. As can beseen from FIG. 3, the scaffold comprises an open framework of struts andlinks that define a generally tubular body. A cylindrical tube of may beformed into this open framework of struts and links described in FIG. 2.

In FIG. 2, a ring 212 is connected to an adjacent ring by several links234, each of which extends parallel to axis A-A. In this firstembodiment of a scaffold pattern (pattern 200) four links 234 connectthe interior ring 212, which refers to a ring having a ring to its leftand right in FIG. 2, to each of the two adjacent rings. Thus, ring 212 bis connected by four links 234 to ring 212 c and four links 234 to ring212 a. Ring 212 d is an end ring connected to only the ring to its leftin FIG. 2.

A ring 212 is formed by struts 230 connected at crowns 207, 209 and 210.A link 234 is joined with struts 230 at a crown 209 (W-crown) and at acrown 210 (Y-crown). A crown 207 (free-crown) does not have a link 234connected to it. Preferably the struts 230 that extend from a crown 207,209 and 210 at a constant angle from the crown center, i.e., the rings212 are approximately zig-zag in shape, as opposed to sinusoidal forpattern 200, although in other embodiments a ring having curved strutsis contemplated. As such, in this embodiment a ring 212 height, which isthe longitudinal distance between adjacent crowns 207 and 209/210 may bederived from the lengths of the two struts 230 connecting at the crownand a crown angle θ. In some embodiments the angle θ at different crownswill vary, depending on whether a link 234 is connected to a free orunconnected crown, W-crown or Y-crown.

The zig-zag variation of the rings 212 occurs primarily about thecircumference of the scaffold (i.e., along direction B-B in FIG. 2). Thestruts 212 centroidal axes lie primarily at about the same radialdistance from the scaffold's longitudinal axis. Ideally, substantiallyall relative movement among struts forming rings also occurs axially,but not radially, during crimping and deployment. Although, as explainedin greater detail, below, polymer scaffolds often times do not deform inthis manner due to misalignments and/or uneven radial loads beingapplied.

The rings 212 are capable of being collapsed to a smaller diameterduring crimping and expanded to a larger diameter during deployment in avessel. According to one aspect of the disclosure, the pre-crimpdiameter (e.g., the diameter of the axially and radially expanded tubefrom which the scaffold is cut) is always greater than a maximumexpanded scaffold diameter that the delivery balloon can, or is capableof producing when inflated. According to one embodiment, a pre-crimpdiameter is greater than the scaffold expanded diameter, even when thedelivery balloon is hyper-inflated, or inflated beyond its maximum usediameter for the balloon-catheter.

Pattern 200 includes four links 237 (two at each end, only one end shownin FIG. 2) having structure formed to receive a radiopaque material ineach of a pair of transversely-spaced holes formed by the link 237.These links are constructed in such a manner as to avoid interferingwith the folding of struts over the link during crimping, which, asexplained in greater detail below, is necessary for a scaffold capableof being crimped to a diameter of about at most Dmin or for a scaffoldthat when crimped has virtually no space available for a radiopaquemarker-holding structure.

A second embodiment of a crush-recoverable scaffold structure has thepattern 300 illustrated in FIG. 4. Like the pattern 200, the pattern 300includes longitudinally-spaced rings 312 formed by struts 330. A ring312 is connected to an adjacent ring by several links 334, each of whichextends parallel to axis A-A. The description of the structureassociated with rings 212, struts 230, links 234, and crowns 207, 209,210 in connection with FIG. 2, above, also applies to the respectiverings 312, struts 330, links 334 and crowns 307, 309 and 310 of thesecond embodiment, except that in the second embodiment there are onlythree struts 334 connecting each adjacent pair of rings, rather thanfour. Thus, in the second embodiment the ring 312 b is connected to thering 312 c by only three links 334 and to the ring 312 a by only threelinks 334. A link formed to receive a radiopaque marker, similar to link237, may be included between 312 c and ring 312 d.

FIGS. 5A and 5B depict aspects of the repeating pattern of closed cellelements associated with each of the patterns 300 and 200, respectively.FIG. 5A shows the portion of pattern 300 bounded by the phantom box VAand FIG. 5B shows the portion of pattern 200 bounded by the phantom boxVB. Therein are shown cell 304 and cell 204, respectively. In FIGS. 5A,5B the vertical axis reference is indicated by the axis B-B and thelongitudinal axis A-A. There are four cells 204 formed by each pair ofrings 212 in pattern 200, e.g., four cells 204 are formed by rings 212 band 212 c and the links 234 connecting this ring pair, another fourcells 204 are formed by rings 212 a and 212 b and the links connectingthis ring pair, etc. In contrast, there are three cells 304 formed by aring pair and their connecting links in pattern 300.

Referring to FIG. 5A, the space 336 and 336 a of cell 304 is bounded bythe longitudinally spaced rings 312 b and 312 c portions shown, and thecircumferentially spaced and parallel links 334 a and 334 c connectingrings 312 b and 312 c. Links 334 b and 334 d connect the cell 304 to theright and left adjacent ring in FIG. 3, respectively. Link 334 bconnects to cell 304 at a W-crown 309. Link 334 d connects to cell 304at a Y-crown 310. A “W-crown” refers to a crown where the angleextending between a strut 330 and the link 336 at the crown 310 is anobtuse angle (greater than 90 degrees). A “Y-crown” refers to a crownwhere the angle extending between a strut 330 and the link 336 at thecrown 309 is an acute angle (less than 90 degrees). The same definitionsfor Y-crown and W-crown also apply to the cell 204. There are eightconnected or free crowns 307 for cell 304, which may be understood aseight crowns devoid of a link 334 connected at the crown. There are oneor three free crowns between a Y-crown and W-crown for the cell 304.

Additional aspects of the cell 304 of FIG. 5A include angles for therespective crowns 307, 309 and 310. Those angles, which are in generalnot equal to each other (see e.g., FIG. 6A for the “V2” and “V23”embodiments of scaffold having the pattern 300), are identified in FIG.5A as angles 366, 367 and 368, respectively associated with crowns 307,309 and 310. For the scaffold having the pattern 300 the struts 330 havestrut widths 361 and strut lengths 364, the crowns 307, 309, 310 havecrown widths 362, and the links 334 have link widths 363. Each of therings 312 has a ring height 365. The radii at the crowns are, ingeneral, not equal to each other. The radii of the crowns are identifiedin FIG. 5A as radii 369, 370, 371, 372, 373 and 374.

Cell 304 may be thought of as a W-V closed cell element. The “V” portionrefers to the shaded area 336 a that resembles the letter “V” in FIG.6A. The remaining un-shaded portion 336, i.e., the “W” portion,resembles the letter “W”.

Referring to FIG. 5B, the space 236 of cell 204 is bounded by theportions of longitudinally spaced rings 212 b and 212 c as shown, andthe circumferentially spaced and parallel links 234 a and 234 cconnecting these rings. Links 234 b and 234 d connect the cell 204 tothe right and left adjacent rings in FIG. 2, respectively. Link 234 bconnects to cell 236 at a W-crown 209. Link 234 d connects to cell 236at a Y-crown 210. There are four crowns 207 for cell 204, which may beunderstood as four crowns devoid of a link 234 connected at the crown.There is only one free crown between each Y-crown and W-crown for thecell 204.

Additional aspects of the cell 204 of FIG. 5B include angles for therespective crowns 207, 209 and 210. Those angles, which are in generalnot equal to each other (see e.g., FIG. 6B for the “V59” embodiment of ascaffold having the pattern 200), are identified in FIG. 5B as angles267, 269 and 268, respectively associated with crowns 207, 209 and 210.For the scaffold having the pattern 200 the struts 230 have strut widths261 and strut lengths 266, the crowns 207, 209, 210 have crown widths270, and the links 234 have link widths 261. Each of the rings 212 has aring height 265. The radii at the crowns are, in general, not equal toeach other. The radii of the crowns are identified in FIG. 5A as innerradii 262 and outer radii 263.

Cell 204 may be thought of as a W closed cell element. The space 236bounded by the cell 204 resembles the letter “W”.

Comparing FIG. 5A to FIG. 5B one can appreciate that the W cell 204 issymmetric about the axes B-B and A-A whereas the W-V cell 304 isasymmetric about both of these axes. The W cell 204 is characterized ashaving no more than one crown 207 between links 234. Thus, a Y-crowncrown or W-crown is always between each crown 207 for each closed cellof pattern 200. In this sense, pattern 200 may be understood as havingrepeating closed cell patterns, each having no more than one crown thatis not supported by a link 234. In contrast, the W-V cell 304 has threeunsupported crowns 307 between a W-crown and a Y-crown. As can beappreciated from FIG. 5A, there are three unsupported crowns 307 to theleft of link 334 d and three unsupported crowns 307 to the right of link334 b.

A third embodiment of a crush-recoverable scaffold structure (referredto herein as “V79”) has the pattern 400 illustrated in FIG. 7. Pattern400 is an image of a portion of a scaffold post deployment which wasgenerated the Visicon Finescan™ Stent Inspection System, and isdiscussed in detail in the Examples. Therefore, there is apparent acertain amount of distortion in the angles between the various struts.Like the patterns 200 and 300, the pattern 400 includeslongitudinally-spaced rings 412 formed by struts 430. A ring 412 isconnected to an adjacent ring by links 434, each of which extendsparallel to axis A-A. The description of the structure associated withrings 212, struts 230, links 234, and crowns 207, 209, 210 in connectionwith FIG. 2, above, also applies to the respective rings 412, struts430, links 434 and crowns 407, 409 and 410 of the third embodiment,except that in the third embodiment there are only two struts 434connecting each adjacent pair of rings, rather than four. Thus, in thethird embodiment the ring 412 b is connected to the ring 412 c by onlytwo links 434 and to the ring 412 a by only two links 434.

The sequence of crests starting at a W-crown and going around acircumference of a ring is: W-crown, 3-free crowns, Y-crown, 3 freecrowns, W-crown, etc. The pattern shown in FIG. 7 has 2 W-crowns and 2-Ycrowns. Thus, there are either 2 or 3 free crowns between a W-crown andY-crown. A link formed to receive a radiopaque marker, similar to link237, may be included between two rings.

FIG. 8 shows the portion of pattern 400 bounded by the phantom box VC.Therein are shown cell 404. In FIG. 8, the vertical axis reference isindicated by the axis B-B and the longitudinal axis A-A. There are twocells 404 formed by each pair of rings 412 in pattern 400, e.g., twocells 404 are formed by rings 412 b and 412 c and the links 434connecting this ring pair, another two cells 404 are formed by rings 412a and 412 b and the links connecting this ring pair, etc.

Referring to FIG. 8, the space 436 and 436 a of cell 404 is bounded bythe longitudinally spaced rings 412 b and 412 c portions shown, and thecircumferentially spaced and parallel links 434 a and 434 c connectingrings 412 b and 412 c. Links 434 b and 434 d connect the cell 404 to theright and left adjacent ring in FIG. 7, respectively. Link 434 bconnects to cell 404 at a W-crown 409. Link 434 d connects to cell 404at a Y-crown 410. There are 12 unconnected or free crowns 407 for cell404, which may be understood as 12 crowns devoid of a link 434 connectedat the crown. There are three free crowns between a Y-crown and W-crownfor the cell 404.

Additional aspects of the cell 404 of FIG. 8 include angles for therespective crowns 407, 409 and 410. Those angles, which are in generalnot equal to each other (see e.g., FIG. 6A for the “V2” and “V23”embodiments of scaffold having the pattern 300), are identified in FIG.8 as angles 466, 467 and 468, respectively associated with crowns 407,409 and 410. For the scaffold having the pattern 400 the struts 430 havestrut widths 461 and strut lengths 464, the crowns 407, 409, 410 havecrown widths 462, and the links 434 have link widths 463. Each of therings 412 has a ring height 465. The radii at the crowns are, ingeneral, not equal to each other. The radii of the crowns are identifiedin FIG. 8 as radii 469, 470, 471, 472, 473 and 474.

Cell 404 may be thought of as a W-W closed cell element. The unshadedarea 436 and unshaded area 436 a each resemble the letter “W”.

As described above, crush recoverable copolymer scaffolds also exhibitself-expansion or shape memory properties when in a physiologicalenvironment. The copolymer crush-recoverable scaffolds may self expandafter expansion to a target diameter. The deployed scaffold may have aresidual outward force applied the vessel wall due to the shape memoryproperties. The self expanding properties result in no or reduced recoilafter expansion to the target diameter. The recoil may be less than 1 or2%.

The crush recoverable copolymer scaffolds have greater fractureresistance than a PLLA scaffold. In particular, the inventors have foundthat copolymer scaffolds have few or no fractures when expanded from acrimped diameter to an expanded target diameter, for example, 8 to 10mm. In addition, the crush recoverable copolymer scaffolds would havefewer fractured or broken struts during a time after implantation basedon axial and bending fatigue data for such scaffolds.

As indicated above, the rubbery polymer of the copolymer has a tendencyto decrease the radial strength and radial stiffness of the scaffolds.The composition of the rubbery polymer in the copolymer of the crushrecoverable scaffold is such that the scaffold has the desired radialstrength and radial stiffness while also exhibiting the shape memoryproperties with reduced or no recoil from the target diameter, desiredcrush recoverability, and shape memory properties. The composition thecopolymer of the crush recoverable scaffolds may be 0.1 to 10 mol %, 1to 10 mol %, 0.1 to 5 mol %, 3 to 5 mol %, 5 to 10 mol %, or 8 to 10 mol%.

The stent scaffolds may be formed by extruding polymer tubes made of thecopolymer and laser cutting the tubes to form a scaffold. Thefabrication methods of a bioabsorbable stents described herein caninclude the following steps:

(1) forming a polymeric tube using extrusion,

(2) radially deforming the formed tube,

(3) forming a stent scaffolding from the deformed tube by lasermachining a stent pattern in the deformed tube with laser cutting,

(4) optionally forming a therapeutic coating over the scaffolding,

(5) crimping the stent over a delivery balloon, and

(6) sterilization with election-beam (E-Beam) radiation.

In the extrusion step, a polymer is processed in an extruder above themelting temperature of the copolymer.

In step (2) above, the extruded tube may be radially deformed toincrease the radial strength of the tube, and thus, the finished stent.The increase in strength reduces the thickness of the struts required tosupport a lumen with the scaffold when expanded at an implant site. Inexemplary embodiments, the strut thickness can be 100-200 microns, ormore narrowly, 120-180, 140-160, or 160-200 microns.

Detailed discussion of the manufacturing process of a bioabsorbablestent can be found elsewhere, e.g., U.S. Patent Publication Nos.20070283552 and 20120073733 which are incorporated by reference herein.

“Molecular weight refers to either number average molecular weight (Mn)or weight average molecular weight (Mw). References to molecular weightherein refer to either Mn or Mw, unless otherwise specified.

“Semi-crystalline polymer” refers to a polymer that has or can haveregions of crystalline molecular structure and amorphous regions. Thecrystalline regions may be referred to as crystallites or spheruliteswhich can be dispersed or embedded within amorphous regions.

The “glass transition temperature,” Tg, is the temperature at which theamorphous domains of a polymer change from a brittle vitreous state to asolid deformable or ductile state at atmospheric pressure. In otherwords, the Tg corresponds to the temperature where the onset ofsegmental motion in the chains of the polymer occurs. When an amorphousor semi-crystalline polymer is exposed to an increasing temperature, thecoefficient of expansion and the heat capacity of the polymer bothincrease as the temperature is raised, indicating increased molecularmotion. As the temperature is increased, the heat capacity increases.The increasing heat capacity corresponds to an increase in heatdissipation through movement. Tg of a given polymer can be dependent onthe heating rate and can be influenced by the thermal history of thepolymer as well as its degree of crystallinity. Furthermore, thechemical structure of the polymer heavily influences the glasstransition by affecting mobility.

The Tg can be determined as the approximate midpoint of a temperaturerange over which the glass transition takes place. [ASTM D883-90]. Themost frequently used definition of Tg uses the energy release on heatingin differential scanning calorimetry (DSC). As used herein, the Tgrefers to a glass transition temperature as measured by differentialscanning calorimetry (DSC) at a 20° C./min heating rate.

The “melting temperature” (Tm) is the temperature at which a materialchanges from solid to liquid state. In polymers, Tm is the peaktemperature at which a semicrystalline phase melts into an amorphousstate. Such a melting process usually takes place within a relativenarrow range (<20° C.), thus it is acceptable to report Tm as a singlevalue.

“Stress” refers to force per unit area, as in the force acting through asmall area within a plane. Stress can be divided into components, normaland parallel to the plane, called normal stress and shear stress,respectively. Tensile stress, for example, is a normal component ofstress applied that leads to expansion (increase in length). Inaddition, compressive stress is a normal component of stress applied tomaterials resulting in their compaction (decrease in length). Stress mayresult in deformation of a material, which refers to a change in length.“Expansion” or “compression” may be defined as the increase or decreasein length of a sample of material when the sample is subjected tostress.

“Strain” refers to the amount of expansion or compression that occurs ina material at a given stress or load. Strain may be expressed as afraction or percentage of the original length, i.e., the change inlength divided by the original length. Strain, therefore, is positivefor expansion and negative for compression.

“Strength” refers to the maximum stress along an axis which a materialwill withstand prior to fracture. The ultimate strength is calculatedfrom the maximum load applied during the test divided by the originalcross-sectional area.

“Modulus” and “stiffness” may be defined as the ratio of a component ofstress or force per unit area applied to a material divided by thestrain along an axis of applied force that results from the appliedforce. The modulus or the stiffness typically is the initial slope of astress-strain curve at low strain in the linear region. For example, amaterial has both a tensile and a compressive modulus.

The tensile stress on a material may be increased until it reaches a“tensile strength” which refers to the maximum tensile stress which amaterial will withstand prior to fracture. The ultimate tensile strengthis calculated from the maximum load applied during a test divided by theoriginal cross-sectional area. Similarly, “compressive strength” is thecapacity of a material to withstand axially directed pushing forces.When the limit of compressive strength is reached, a material iscrushed.

“Toughness” is the amount of energy absorbed prior to fracture, orequivalently, the amount of work required to fracture a material. Onemeasure of toughness is the area under a stress-strain curve from zerostrain to the strain at fracture. The units of toughness in this caseare in energy per unit volume of material. See, e.g., L. H. Van Vlack,“Elements of Materials Science and Engineering,” pp. 270-271,Addison-Wesley (Reading, P A, 1989).

EXAMPLES

Scaffold samples were prepared with two poly(L-lactide-co-caprolactone)molar compositions of PLLA/PCL, 90:10 and 90:5. For comparison, PLLAscaffold samples were also prepared. The samples were tested aftere-beam sterilization. Scaffolds with a poly(DL-lactide)/Everolimuscoating were also tested. All scaffold samples were the V79 pattern.

Processing:

90:10 PLLA-PCL with intrinsic viscosity (IV) of 3.2 was processed byextrusion (water content 310 ppm, temp approximately 380 deg F.) to forma tube with 0.059 in ID. The tube was expanded to 8 mm OD at atemperature above its Tg but lower than Tm. It was then laser cut into aV79 scaffold pattern. L-lactide monomer content in extruded tubing andexpanded tubing was 0.12%.

95:5 PLLA-PCL with IV of 3.8 was processed by extrusion (water content320 ppm, temp approximately 370 deg F.) to form a tube with 0.051 in ID.The tube was then expanded to 7 mm OD at a temperature above its Tg butlower than Tm. It was then laser cut into a V79 scaffold pattern.L-lactide monomer content in extruded tubing and expanded tubing was0.11%.

The scaffold samples were crimped onto a folded Omnilink Elite PTAballoons (Abbott Vascular Inc., Santa Clara, Calif.) at 48 deg C.,packaged, and sterilized by e-beam.

Molecular Weight:

The Mn of the 90:10 expanded tubing sample was approximately 150 kDa andthe Mn and of sterilized scaffold was approximately 68 kDa. The Mn ofthe 95:5 expanded tubing sample was approximately 100 kDa and the Mn andof sterilized scaffolds was 82 kDa.

Scaffold Dislodgement Force:

A scaffold dislodgement force for a scaffold crimped to a balloon refersto the maximum force applied to the scaffold along its longitudinal axisthat the scaffold is able to resist before dislodging from the balloon.Scaffold dislodgement force for the samples was measured by a tape test.The scaffold dislodgement force for the 90:10 samples was 1.34±0.35 lband the 95:5 samples was 1.00±0.10 lb.

Radial Strength and Stiffness

Radial strength and stiffness were measured on an MSI tester (MachinesSolutions Inc., Flagstaff, Ariz.) and radial strength for eachcomposition was greater than 0.4 N/mm. FIG. 9 shows the radial strengthand stiffness for PLLA scaffolds, 95:5 PLLA-PCL uncoated scaffolds, 95:5PLLA-PCL coated scaffolds, 90:10 PLLA-PCL uncoated scaffolds, and 90:10PLLA-PCL uncoated scaffolds.

Post-Dilation to Fraction:

The scaffold samples were dilated from the crimped diameter and theminimum OD at fracture was observed. As shown in Table 2, both 95:5 and90:10 scaffolds demonstrated a significantly higher fracture diameterthan the PLLA scaffold theoretical fracture diameter of 9.6 mm.

TABLE 2 Post-Dilation to fracture of scaffolds. Scaffold TypePost-Dilation to Fracture (mm) PLLA 9.6 mm (theoretical) 95:5 PLLA-PCL10.0 +/− 0.1 mm. 90:10 PLLA-PCL 10.7 +/− 0.8 mm.

Recoil:

The inward and outward recoil of the 90:10 PLLA-PCL scaffold wasstudied. The scaffold was balloon expanded to 6 mm, the balloon wasdeflated and then the scaffold diameter was recorded with time. FIG. 10shows the diameter of the scaffold as a function of time afterexpansion. The diameter recoiled inward for about the first 60 minfollowed by recoil outward with a significant OD growth as a function oftime at least up to 28 days.

Crush Recovery:

The crush recovery of the PLLA and 90:10 PLLA-PCL scaffolds were tested.FIG. 11 shows the results for both of the scaffolds.

Axial Fatigue:

The PLLA, 95:5 PLLA-PCL, and 90:10 PLLA-PCL scaffolds were subjected toaxial fatigue testing.

The scaffolds coated with silicone sealant were deployed inside siliconetubing with two different hardnesses. The samples were dried for 24hours and the tubes were stretched axially by 7%. The scaffold sampleswere then subjected to 500k cycles at 1 Hz in circulating water at 37°C. The scaffold samples were inspected for fractures at different timepoints.

The results demonstrated the PLLA-PCL scaffolds had reduced scaffolddiscontinuities at 500K cycles compared to PLLA. FIG. 12 showsdiscontinuity count in ring and connector links for each scaffold. FIG.13 shows discontinuity percentage in ring and connector links for eachscaffold.

Imaging of Scaffolds Post Deployment

Images of the scaffolds post deployment were generated the VisiconFinescan™ Stent Inspection System, Visicon Inspection Technologies, LLC(Napa, Calif.). The system employs a scan camera to generate a flat,unrolled view of a scaffold. In operation, the scaffold is mounted on amandrel with a fine diffuse surface. This mandrel is held under thelinear array camera and rotated by the system electronics and is used totrigger the camera to collect a line of image data in a preciseline-by-line manner.

FIGS. 14 and 15 depict Finescan images of the 95:5 and the 90:10PLLA-PCL scaffolds, respectively, post deployment. The 95:5 scaffold hasminimal disruption in its structure.

While particular embodiments of the present invention have been shownand described, it will be obvious to those skilled in the art thatchanges and modifications can be made without departing from thisinvention in its broader aspects. Therefore, the appended claims are toencompass within their scope all such changes and modifications as fallwithin the true spirit and scope of this invention.

What is claimed is:
 1. A stent comprising a scaffold formed from apolymer tube configured for being crimped to a balloon, the scaffoldhaving a pattern of interconnected struts and the scaffold having anexpanded diameter when expanded from a crimped state by the balloon,wherein the scaffold attains greater than about 80% of its diameterafter being crushed to at least 50% of its expanded diameter; whereinthe scaffold has a radial stiffness greater than 0.3 N/mm and whereinthe scaffold is made from a shape memory random copolymer ofpoly(L-lactide) (PLLA) and a rubbery polymer that is 0.1 to 10 wt % ormol % of the copolymer, wherein the scaffold exhibits self expandingproperties at 37 deg C. in physiological conditions.